Magnetic Resonance Imaging (MRI) is a well-known procedure for obtaining detailed, two and three-dimensional images of patients based on nuclear magnetic resonance (NMR) principles. MRI is well suited for the imaging of soft tissues and is primarily used for diagnosing internal injuries.
Typical MRI systems include a magnet capable of producing an intense, homogenous magnetic field around a patient or portion of the patient; a radio frequency (RF) transmitter and receiver system, including a receiver coil also surrounds a portion of the patient; a magnetic gradient system localizes a portion of the patient; and a computer processing/imaging system, which receives the demodulated signals from the receiver system and processes the signals into interpretable data, such as visual images.
The superconducting magnet is used in conjunction with a magnetic gradient coil assembly, which is sequentially pulsed to create a sequence of controlled gradients in the main magnetic field during an MRI data gathering sequence. The superconducting magnet and the magnetic gradient coil assembly include the radio frequency (RF) coil on an inner circumferential side of the magnetic gradient coil assembly. The controlled sequential gradients are. effectuated throughout a patient imaging volume (patient bore) which is coupled to at least one MRI (RF) coil or antenna. The RF coils and an RF shield are typically located between the magnetic gradient coil assembly and the patient bore.
As a part of a typical MRI, RF signals of suitable frequencies are transmitted into the patient bore. Nuclear magnetic resonance (NMR) responsive RF signals are received from the patient via the RF coils. Information encoded within the frequency and phase parameters of the received RF signals, by the use of an RF circuit, is processed to form visual images. These visual images represent the distribution of NMR nuclei within a cross-section or volume of the patient, within the patient bore.
In modem MRI, the demand for high spatial and temporal resolution necessitates the use of high static magnetic field. Active electric coils are used to drive spatial gradients into the static magnetic field. Enhanced imaging sequences typically demand high amplitude gradient fields, rapid field transitions, and large duty cycles in order to improve resolution and scan time unfortunately, these properties also drive the power dissipation higher and thus cause higher temperatures in the scanner.
At high frequencies, the field generated by the volume radio frequency coils for different patient sections, e.g. head and body, becomes inhomogeneous as a result of electrical properties of patient tissue. Homogeneity becomes important during MR transmit pulses for uniform distribution of flip angles, which in turn are required for homogenous contrast or fat saturation. At these high frequencies a further problem is that a high amount of energy may be dissipated in the tissue of the patient, which is generally undesirable.
The primary design criterion for RF transmit coils for MRI is a uniform transverse RF magnetic field, which is referred to herein as B1. Currently, most high frequency transmit coils are designed as periodic, symmetric structures that are resonant at the imaging frequency, as determined by the static magnetic field (B0) strength. These coils are excited by one or more voltage sources.
The distribution of currents on the coil elements/axial conductors or rungs is determined by the symmetry of the coil structure. At field strengths of 3T and above, electric properties such as the dielectric constant and conductivity of the load lead to B1 inhomogeneity due to wavelength effects and perturbation of the coil current distribution from the ideal.
The B1 homogeneity under such conditions may be optimized by adjusting the amplitudes and phases of the currents on the rungs. However, such adjustments require independent control of current amplitudes and phases on each rung of the resonant coil. Due to strong coupling between the rungs of a resonant coil and sensitivity to loading, such independent control has not been possible and homogeneity optimization involves a time consuming and impractical iterative procedure in the absence of exact knowledge of interactions among coil elements and between the coil and load.
It would, therefore, be highly desirable to have a magnetic resonance imaging RF coil with independent control of current phase and amplitude on each conductor, which will improve control over magnetic field homogeneity. It would also be desirable only to selectively excite regions of interest rather than all regions on a patient during a scan. The present invention is directed to these ends.